Known gamma cameras comprise detection devices consisting mainly of a detector, a collimator and data processing means.
The detector can comprise a scintillator material, such as cesium iodide, for instance CsI(Tl), sodium iodide, for instance NaI(Tl), lanthanum bromide (LaBr3) or bismuth germanate (BGO), in association with photo detectors, for instance a matrix of photo diodes, specifically avalanche photo diodes, a CCD matrix or a CMOS sensor. They are called scintillating detectors. The thickness of the scintillating material is in general between a few μm and a few mm. When a photon penetrates in the scintillator material and interacts with it, it produces photons of lower energy, in general in the visible spectrum. These photons are then collected by at least one photo detector coupled to the scintillator material, then transformed in an exploitable electrical signal. A scintillating detector comprises usually a plurality of pixels, each pixel corresponding with at least one photo diode, or at least one pixel of a CCD or CMOS matrix.
Alternatively, the detector comprises at least a semi-conductor detector material, susceptible of being polarized by a cathode and an anode. These electrodes are in general arranged on two opposite faces of the semi-conductor material. This type of detector is called a semi-conductor detector. When a photon penetrates in the semi-conductor material and interacts with it, all or part of its energy is transferred to charge carriers in the semi-conductor material. Since the detector is polarized, the charge carriers migrate towards the electrodes (the anode). They produce then an electrical signal at the terminals of the electrodes. The energy of this electrical signal is proportional with the energy deposited by the photon during the interaction. The electrical signal is collected and then processed. Depending on the nature of the detector, the signal is collected only at the anode (most frequent case), only at the cathode, or at both electrodes. A semi-conductor detector comprises usually a plurality of physical pixels, and each physical pixel corresponds with a charge collection circuit of one electrode.
The collimator allows for a selection of the photons arriving at the detector. It is formed by channels delimited by fine walls called septas. These channels (or more precisely the corresponding holes) can have a circular, hexagonal or square section; they can be parallel, divergent or convergent.
Known gamma cameras can be used either in planar mode, or in tomographic mode. In planar mode, acquisition takes place according to only one angle of view, the head of the camera remains fixed during the whole examination. The projection and therefore the reconstructed image have the inconvenience of not integrating any information about the location and depth of the radioactive elements distributed in the patient. In tomographic mode, the head of the camera, comprising the detector and the collimator, describes a circular or elliptic orbit around the patient and a plurality of projections are acquired at different view angles. Reconstruction techniques are then employed in order to create images that can be interpreted by the practitioner. With the tomographic mode, specific information can be obtained about location and depth of the radioactive elements distributed in the patient.
The performance of detection devices is usually characterized by a certain number of parameters, among which:                1) the spatial resolution, which corresponds with the minimum distance between the two point sources that can be discerned on a planar image (image resulting from projection under only one angle of view) or on a reconstructed image (starting from several projections). Spatial resolution is usually given by the width at mid-height of the punctual or linear dispersion function (obtained by imaging a punctual or linear source, respectively). The spatial resolution is the result of the intrinsic spatial resolution of the detector and the geometric spatial resolution of the collimator;        2) the energy resolution, which illustrates the capacity of a detection device to precisely select photons in function of their energy. It is expressed in percent and is given by the width at mid-height ΔE of the energy response of the device for the energy emission E of the source (in other words of the used radio isotope);        3) the sensitivity, also called effectiveness. Sensitivity can be defined by the relationship between the number of primary photons detected (photons that have not undergone any interaction before reaching the detector material) and the total number of photons emitted by the source over 4π steradians. The sensitivity depends on the effectiveness of the detector (arresting power of the detector material) and the geometric effectiveness of the collimator (which is low because the collimator imposes a strong spatial selection of photons and, because of this, absorbs the majority of them). The lower the sensitivity of a detection device, the longer the acquisition time must be in order to obtain satisfactory statistics about the acquired projections.        
As a general rule, these parameters vary inversely: the improvement of one parameter leads to the degradation of one or more other parameters. This is for instance the case of the spatial resolution and the sensitivity. Similarly, as disclosed by publication “Study of high-energy multihole collimators” (Andreas R. Formiconi, XP011077435), by moving the collimator away from the detector, artifacts caused by the rows of collimator channels can be minimized, but the spatial resolution will be degraded. When designing a detection device, a person skilled in the art must therefore select (specifically in function of the destination of the device) the parameter(s) that he wishes to favor to the detriment of others, or accept a compromise.
Presently, nuclear medicine departments are using mainly cameras called Anger cameras, with a detector comprising a NaI(Tl) scintillator and the collimator formed by parallel channels with hexagonal (honeycomb) section. Such device imposes a compromise between spatial resolution and sensitivity: for a source of 140 keV placed at 10 cm of the collimator, the spatial resolution is 10 mm and the sensitivity, as defined previously, is 10−4. The spatial resolution is particularly limited by the resolution of the scintillator, which is a few mm, for instance 3 mm. In addition, the image has low contrast due to mediocre energy resolution of the scintillator (10% at 140 keV).
The appearance of semi-conductor detectors, such as detectors on the basis of CdTe, CdZnTe (CZT), silicon (Si) or HgI2, has given birth to a new generation of gamma cameras. As opposed to scintillators, semi-conductors are materials with direct conversion: following the interaction of gamma rays, they create, without intermediate stage, electrical charges, in other words electron-hole pairs (electrons migrating towards the anode, the holes towards the cathode). In general, these detectors allow for the use of semi-conductor materials with thickness between a few μm and a few mm, and what is more, without the increase in thickness being accompanied by a drastic degradation of the spatial resolution.
For instance, a prototype known under the name PEGASE integrates a CZT semi-conductor detector and a collimator with parallel channels with square section, in which one hole of the collimator corresponds with one pixel of the detector. The used detector has improved energy resolution (between 3% and 5% at 140 keV, against 10% for the NaI(Tl) scintillator, which translates into a significant increase of image contrast. On the other hand, although its detector possesses better intrinsic spatial resolution, which can go for instance from 1.6 to 2.5 mm, in terms of spatial resolution and sensitivity the PEGASE prototype offers performance which is almost identical to that of Anger gamma camera.
The problem consists therefore in improving the performance of known detection devices, relative to the flux of captured photons, in other words the sensitivity (limited due to the presence of the collimator), and the spatial resolution, in order to finally break the compromise imposed since the birth of the Anger gamma camera.
To this end, the researchers have tried mainly to refine the localization, inside each pixel, of the photon interaction taking place in this pixel. As illustrated in attached FIGS. 1a and 1b, better localization of the interaction allows indeed for a reduction of the angular sector from which the detected photon originates, and consequently an improvement of the localization of the emission source.
A first method has consisted in compensating the too approximate knowledge of the place of interaction in the detector by translations of the collimator or of the detector along the X and Y directions of the detection plane. Throughout the description, the expression “detection plane” designates the frontal plane of the detector situated on the side of the collimator. The creation of a plurality of projections, corresponding to different relative positions of collimator and detector, results then, by combining the different images obtained, in an improvement of the spatial resolution. This first method was proposed with a collimator with holes having a transversal dimension (diameter, side . . . ) smaller than half the dimension of the pixels, and with a high efficiency collimator (and therefore with “large” holes). This first method is however not without inconvenient. It requires the detection device to be equipped with translations means in the detection plane according to two directions, means which are in general complex and costly. By multiplying the number of necessary projections (and or acquisitions), the acquisition time is furthermore increased, which penalizes the consultation cadence of medical imaging centers.
A second method has consisted in diminishing the size of the pixels in order to obtain a degree of pixelation greater than 1 in the detection plane. This is called over-pixelation. Throughout the description, the expression “degree of pixelation in the detection plane” designates the number of pixels extending opposite a collimator channel, according to a direction Z orthogonal to the detection plane. When this degree is greater than 1, each channel of the collimator corresponds with more than one pixel. In other terms, the transversal dimensions of the pixels are smaller than those of the collimator channels. To be noted that under “transversal dimensions of the pixels” is understood the dimensions of each pixel in the detection plane; the expression “transversal dimensions of the channels” designates the dimensions of the hole of each channel in a frontal plane of the collimator, called interior plane, which extends opposite the detection plane.
Such over-pixelation in the detection plane allows for elimination of the translations of collimator or detector in the (X, Y) plane, which were necessary in the first method, while limiting the acquisition time. Over-pixelation can be obtained physically or virtually.
We speak of physical over-pixelation when the detector presents a plurality of physical pixels opposite each collimator channel. Physical over-pixelation necessitates nevertheless multiplication of the electronic paths and can pose certain technological problems linked to the size of the physical pixels.
In a variant, it was proposed to create virtual over-pixelation in the detection plane, by virtually increasing the number of pixels of the detector. This approach, described in publication “CZT Pixel detectors for improved SPECT imaging” (Guillaume Montémont, XP031419399) allows for dividing each physical pixel in a plurality of virtual pixels in the (X, Y) detection plane. As disclosed by the publication “An approach to sub-pixel spatial resolution in room temperature X-ray detector arrays with good energy resolution” (W. K. Warburton), it is possible to determine the (x, y) coordinates of the barycenter of the cloud of charges produced by an interaction, starting from electrical signals simultaneously generated by this interaction on a plurality of adjacent anodes. It is then assumed that the interaction was detected uniquely by the virtual pixel corresponding with the coordinates of the barycenter. Without this kind of method, only the coordinates are used of the physical pixel that has collected the maximum signal. With barycentric localization, virtual pixel sizes (surfaces) can be achieved today which are one tenth of the size of physical pixels. A collimator channel faces then a plurality of virtual pixels. Compared to physical over-pixelation, virtual over-pixelation has the advantage of being accessible without multiplication of the electronic paths. It is therefore understood that over-pixelation in the detection plane corresponds with multiplication of pixels opposite a collimator channel. This over-pixelation in the detection plane can be physical or virtual, although a combination of physical over-pixelation and virtual over-pixelation can also be envisaged.
Finally, WO2008/046971 proposes to refine even more the localization of the interactions, not only by dividing each pixel in several virtual pixels in the (X, Y) detection plane, but in addition by virtually dividing the thickness of the detector material in a plurality of layers (of 1 mm thickness for instance). In other words, it involves here determining an interaction coordinate according to the Z axis. The detector becomes connected to a 3D matrix of virtual unitary detection elements, called voxels. Here also, the supplied supplementary information allows us to refine the spatial localization of the source and therefore to ameliorate the spatial resolution of the detection device.
The recent evolutions mentioned above have therefore led to detection devices with increased spatial resolution. The progress made remains however unsatisfactory considering that none of the devices proposed until now is fully using the capabilities of semi-conductors. Above all, the sensitivity of these devices remains low.